The present invention relates generally to Positron Emission Tomography (“PET”) scanners, and more particularly to a method of PET scanning and a PET scanner apparatus having increased image resolution.
Positron Emission Tomography (“PET”) is an imaging technique that provides three-dimensional tomographic images of a distribution of positron-emitting isotopes within an object. The object is usually a living human or animal, and the images provide a visual depiction of tissue differences within different portions of the object. A PET procedure involves the introduction of radiolabeled tracing pharmaceuticals into the object, usually through injection or inhalation. The type of pharmaceuticals depends on the function of the tissue under investigation. As the radiolabeled tracing element in the pharmaceutical decays, it produces positrons. Each positron collides with surrounding matter in the sample object before combining with an electron in the sample object. The combination of each positron with an electron in the sample object annihilates both particles, producing a pair of gamma-ray photons. The gamma-ray photons travel away from the annihilation event in opposite directions. If a pair of opposing gamma-ray detectors each detect one of the two gamma-ray photons produced in the annihilation event within a predetermined period of time, usually 5 to 50 nanoseconds, a “coincidence event” is recorded, and it is assumed that the annihilation event producing the gamma-ray photons lies along a straight line between the two detectors.
Conventionally, a PET scanner consists of arrays of gamma-ray detectors, arranged either continuously as one or multiple rings, or as two or more detector plates. Lines of response (“LOR”) are formed between opposing detector pairs in each array. The PET scanner obtains the radioactivity distribution information within the object by detecting annihilation events originating along each LOR. Commercially available PET scanners having one or more rings of detectors are available for PET scanning animal and human subjects. The inside diameters of the one or more rings of detectors for PET scanning animal and human subjects are approximately 20 centimeters and 80 centimeters, respectively. Conventionally, for scanners having detectors arranged in a ring(s), the object is placed at the center of the ring(s) where the sampling is believed to be the highest, hence achieving the best resolution and image quality currently available. For scanners having detector plates, the detector plates rotate around the object to collect data from all angles in order to form a complete set of projections of the distribution. For the same reason as in the ring configuration, the object is conventionally centered between the detector plates. The detection of a large number of annihilation events allows a computer to construct a three-dimensional image of the distribution of radiolabeled pharmaceuticals within the object, which provides valuable information on the kinetics of the pharmaceuticals and functions of the living object.
With these conventional designs, the image spatial resolution of a PET system is determined by several factors, including intrinsic detector spatial resolution, acolinearity of the annihilating gamma ray photons, and positron range of the radioisotopes in the tracing pharmaceuticals. Of these three factors, the last two depend on the type of radioisotopes used and are independent of the scanner design. Therefore, PET scanner manufacturers have been trying to improve scanner spatial resolution by designing new detectors that improve the detector's intrinsic spatial resolution. This is particularly important for very high resolution PET scanners dedicated to small animal imaging, which have become a very powerful tool for the advancement of molecular imaging.
For most animal PET scanners and some state-of-the-art human scanners, discrete scintillation crystals coupled to photodetectors have been used to achieve the highest spatial resolution heretofore possible. For a PET scanner using discrete crystals, the detector intrinsic spatial resolution can not be better than one half of the crystal width. For a PET scanner with ring geometry, the detector pairs form sampling lines with an average sampling distance of half a crystal width. Based on the Nyquist theorem in sampling theory, the smallest object (i.e., the highest frequency of signal) that one system can resolve is twice the size of the sampling distance (i.e., half of the sampling frequency). In order to achieve image spatial resolution that approaches the theoretical limit, where the detector intrinsic spatial resolution equals one half the crystal width, conventional PET scanners require smaller sampling distances. Many attempts have been made to increase the sampling resolution. For example, certain designs move the detector or the object by a fraction of the detector width. Other designs stack discrete crystals in multiple offset layers. With these designs, image resolution can begin to approach the detector intrinsic resolution. However, conventional PET scanners have been unable to achieve image resolution higher than the detector intrinsic spatial resolution regardless of the type of gamma-ray detector employed. This is true for PET scanners with scintillation detectors, ionization chambers, semiconductor detectors and other types of gamma-ray detectors.
Several techniques have been developed in other imaging arts to resolve structures smaller than the detector intrinsic spatial resolution. One example is a gamma camera coupled to a pinhole collimator that produces a “magnified” image of the object, allowing image resolution of objects smaller than the detector intrinsic spatial resolution. The drawback of this design is a significant reduction of detecting efficiency.
An example of another imaging device is a Compton camera having two detectors placed to one side of a photon source. The detectors of a Compton camera are designed to sequentially detect a photon that interacts with one and then the other detector. The interaction with the first detector is through the Compton effect while the interaction with the second detector is through the photoelectric effect. The sequential detection of a photon enables the Compton camera to trace the photon's path without using mechanical collimators, such as those in a gamma camera. Therefore, a Compton camera has better sensitivity than the conventional gamma camera. The disadvantage of a Compton camera is its reduced resolution compared with a conventional gamma camera. In contrast to apparatus of the present invention, both detectors of the Compton camera are positioned on one side of a photon source and detect a single photon sequentially. In the present invention detectors positioned in opposite sides of a photon source each detect separate photons traveling in opposite direction from the coincidence event. Also, in contrast to the Compton camera, the present invention does not rely on Compton effect interactions to produce images of the object.